1. Sphere of Invention
This invention relates to X-ray engineering and medical diagnostics which deal with techniques of imaging and visual representation of X-ray radiation in the energy range of 5 keV to 200 keV. In particular, this invention may be useful in medicine for the control and monitoring of pathological changes in a living body. This field of medicine, known as radiology, dates back to the beginning of the XX century when German physicist C. Rontgen discovered penetrating radiation named after him.
This invention can also be practicable for dynamic stream preventive examination of patients where the main task is to reveal major pathologies. The invention may also be used in dentistry which requires X-ray examination of the dentofacial area. Another important role the invention may play is in the field of mammalogy for the examination of mammary glands in women.
In addition to medicine, the invention may be useful in defectoscopy and in non-destructive examination systems in various areas of mechanics, such as weld examination in pipelines. The invention is of particular importance for the quality control of fully armed munitions where the technique is possibly the only way of stream quality control.
The invention may be used in the customs control of dimensional cargo in railroad, air and sea transportation.
Such wide area of possible usage of the invention may be explained partly by the need for non-destructive quality control and diagnostics and is an unusual combination of state-of-art materials comprised of scintillators with matrix-type semi-conductor systems for information reading, computer processing and archiving. Undoubtedly, the invention involves advanced technologies.
2. Current State of Technology
The first X-ray apparatuses were made in the 1920s and comprised an X-ray emitting source and a radiation receiver. Since then, no major changes were implemented in X-ray emitting sources: a vacuum device's high-energy accelerated electron beams bombards a metallic anticathode made of tungsten (W) molybdenum (Mo), or sometimes, of copper (Cu). Emerging impact X-ray radiation is filtered to make it monochromatic and then escapes the device through special materials which are penetrable by radiation type (such as beryllium foil). The resulting X-ray beam has a diameter of several millimeters to ten centimeters. Impenetrable structures or bodies requiring examination are placed into this beam.
For a long time, the only way of visually imaging a change in X-ray quantum density in a beam was by a photoemulsion detector based on silver halides. But due to the relatively low density (2-3 g/cm3) of detecting layers and the low sensitivity of silver halides to X-ray radiation, this method required high exposure to X-ray radiation, thus limiting medical usage.
The first technical solution aimed at significantly lowering the exposure dose for X-ray examined patients were X-ray intensifying screens. These screens allowed for a physical shift of quantum energy. A screen was basically a thin layer of a substance—X-ray luminescent material—which radiated when exposed to X-rays. A screen was usually made in the form of a cassette comprising front and rear screens and a photo-sensitive film. The front screen had a thinner layer of X-ray luminescent material and the rear screen was aimed at stopping X-ray radiation almost completely.
For a long time, the main material of X-ray luminescent materials in the intensifying screens was calcium tungstate (CaWO40 which features a high gravitational density and a medium energy conversion efficiency (6.0-8.0%). These parameters were used as a reference for optimal, compounds of X-ray luminescent material. The basic requirements to these compounds were as follows:
average atomic number N in excess of 40 atomic units;
gravitational density in excess of 4.5 g/cm3;
energy efficiency of X-ray luminescent material emission>6%;
back-glow less than 1*10−3 sec;
spectral maximum of radiation λ>400 nm.
Radiology based on direct interaction between X-ray radiation and living tissues (in medical X-ray diagnostics) or between X-ray radiation and parts of complex systems and structures (in X-ray defectoscopy) allowed for the following imaging characteristics in photo-sensitive film or translucent screens:
resolution 1-0.6 mm per pair of lines;
contrast with a ratio between dark or light fields and background below 30%;
imaging of tiny details of 650-800 μm in size;
after-glow period around 1·10−3 sec.
It should be noted that the radiation stress of a patient was excessive even for the state-of-art techniques at the time (1.0 to 10.0 Roentgen units per examination of gastrointestinal tract) [1].
The low characteristics of X-ray diagnostics in 40-70-s necessitated the development of new methods of diagnostics based on other physical principles; thus, in 1964 [2] the first X-ray electron-optical image converting systems (EOIC) were proposed, which managed, the primary conversion (transduction) of X-ray radiation to visible light with further multiple intensification and transformation of the light signal into small picture frame television images [2]. Transducers of the first EOICs were based on halogenide luminescent materials, such as water-soluble cesium iodide, which made the production technology of equipment significantly complex.
At the same time, a method of rapid X-ray photoroentgenography was developed which involved the projection of an image formed on a large translucent luminescent screen onto photographic film by means of a large aperture lens system. This method became convenient for applications where many patients required examination in a short period of time. Photoroentgenography resulted in the discovery of large niduses only.
Beginning in the mid-70-s, the era of rear-earth X-ray luminescent materials began [4], first with primary oxysulfides (Y2O2S:Tb, Gd2O2S:Tb), and then with oxybromide (LaOBr). The main achievements and challenges of this period of material and screen development are summarized below [5].
This period of rapid development of materials for radiography brought important scientific results, in particular, the requirements adjustment to the chemical bonding of an X-ray luminescent matrix, as well as the achievement of good experimental results with visual light energy output efficiency under X-ray or gamma-radiation (e.g. 22% for Y2O2S:Tb which paved the way to a new level of knowledge.
Comparison, in the La2O3—La2O2S—LaOBr:Tb row shows a significant impact in covalent-type bonding in luminescent matrices, which previously had mostly ionic bonding. Research results and certain generalizations [6] may be summarized in the following table of optimum parameters for X-ray luminescent materials.
TABLE 1WavelengthMaximumAtomicEnergyof spectralrange ofAfter-number,Density,efficiency,maximum,radiationglow,CompositionNg/cm3%nmhardness keVmsCaWO461.86.16-9420 80-100 1-1.5ZnCdS:Ag384.814-1856080-901-2CsJ:Tl414.212-18550800.001Y2O2S:Tb364.9521-22383.47860-701-3Gd2O5:Tb59.96.0020-24545100-1201-3LaOBr:Tb49.35.718-20543 80-1101-2
Most notably, this period of radiology development resulted in a 3-4 time decrease of radiation stress on patients, especially in children. Along with this, a significantly higher ratio of X-ray absorption from new X-ray-sensitive materials results in the rejection of traditionally used X-ray screen coarse-grain materials and the utilization of medium-grain materials which increase the resolution of intensifying screens by 20-40%. This was enough to allow the naked eye to see calcified focuses in the mammas of women. This was the beginning of mammography as a preventive field of practical radiography.
At the same time, PHILLIPS proposed using new column X-ray-sensitive coatings (CsJ:Tl) in the screens of EOICs, which had the advantage of not dispersing light due to the light-conducting properties of column microcrystalls of the cesium iodide. The image quality of these apparatuses was as high as in succeeding serial apparatuses which had screens with gadolinium oxysulfide. EOICs allowed for the observation of interaction between soft tissues with X-ray contrast substances, such as barium sulfate or tantalic gadolinium (GdTaO4) (FIG. 1) to reveal ulcerous focuses or other pathologies in a patient's body. The brightness of screens in X-ray devices was improved and reached a threshold of direct registration (brightness level 2-3 cd/m2) by means of an optical image transfer and/or intensifier also based on CCD-matrix [6].
At the same time, it allowed for a decrease in, the energy limit of registration for soft X-ray radiation featuring an energy of E=100-1000 eV. This technology was later used in deep-space apparatuses [6].
The production of highly sensible CCD-matrixes with 10−2 lux of light threshold started the development of advanced digital X-ray-sensitive devices [7] which form images with a brief delay during patient examination.
This new stage of real-time radiography development lasts until today. This stage involves:
improvement of dispersiveness of the most effective material based on Gd2O2S:Tb [8];
selection of compositions based on this luminescent material [9];
creation of X-ray microdetectors [10];
improvement of silicon matrices [11];
creation of the first types of digital X-ray detectors [12];
digital X-ray detector [13];
utilization of white reflecting coatings based on Ta2O5 in [14];
utilization of optically transparent ceramics based on Lu2O3Eu in the new detector [15].
The most recent publication on this topic is an article by Korean scientists [16] who proposed the construction and production technology of a multi-element X-ray sensitive layer of Gd2O2S:Tb in polyethylene press-work having elements coated with reflecting film made of Cr—Al (6000 Å thick). The authors noted a 1.5-2 time decrease in the X-ray luminescent material radiation compared to the solid layer of luminescent material. Yet, the modulation transfer function in the image formed on the structured X-ray sensitive screen is somewhat higher with several extrema of the frequencies close to the geometrical dimensions of screen elements.
Despite some advantages in the detector described in [16] such as a decrease in overall X-ray radiation which falls into a pixelled (multi-element) scintillator only, this construction has several major disadvantages:
reduction of radiation intensity of multi-element screens made of Gd2O2S:Tb;
shallow thickness of X-ray luminescent material allowing for X-ray radiation to reach photodiodes directly causing their degradation;
complex production of microscopic multi-element detectors due to the utilization of the photolithographic process; thus, the original article cites only a small sample of the screen 2 by 3 cm in size;
low contrast of the image on a scintillator; to enlarge the image, the scintillator is additionally covered with a blackening graphite grid;
the shallow thickness of Gd2O2S:Tb luminescent material allows for the use of low accelerating voltages in the X-ray tube only, e.g. 45 keV, which is suitable for limited application only, for example, in dentofacial examination.
These disadvantages were considered in the publication [17] which we used as a prototype for our invention in which the authors suggest a return to radiation sources made of CsJ:Tl with columns of 4-7 μm thick. Elements up to 16 mm high comprised of such structures were used to create a complete scintillator. The authors claimed that such a detector had a modulation transfer function MTF=40% at a resolution of 4 pairs of lines per mm and a MTF=10-20% at a resolution of up to 8 pairs of lines per mm along with an image contrast decrease.
Despite some advantages presented by high quantum detectivity DQL=0.28, the authors suggested that use of CsJ:Tl is not necessarily effective due to defects in microcrystals. With all this, as the authors insist, the intensification of patient radiation stress may be only partly justified by the high definition of the detector.
Despite the various advantages of the prototype detector, such as high quantum detectivity which is singularly prominent at low energies, the prototype had many significant deficiencies. First of all, it has a narrow range of exciting energy of X-ray radiation from 35 keV to 60 keV which is insufficient for a complete medical examination. The second issue is that the radiation load may reach high values of ten roentgens, especially if complex pathologies are to be examined, or if X-ray contrast substance is used. Third, due to the small size of each of the structural elements of the detector (16 mm), the resulting image would be fogged because of the discontinuity of each element.
Fourth, the hidrophylic behavior and temperature-sensitive character of cesium iodide CsJ:Tl requires a comprehensive sealing of detector elements and protection from moisture, which poses a complex problem in view of the elements' tiny size.
Fifth, it should be noted that the production of column structures of CsJ along with interaction with extremely toxic thallium Tl is a very complex and environmentally-prone problem that may be solved by the use of advanced-technology rooms with closed-circuit atmosphere and induced air.
Thus, the series of disadvantages of the existing X-ray detector design, such as the narrow energy operating range, the discontinuity of the imaging field, the low hydrolytic stability and the durability, creates a need for the proposed X-ray detector.